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Impact Loading and Locomotor-RespiratoryCoordination Significantly Influence Breathing Dynamicsin Running HumansMonica A. Daley1*, Dennis M. Bramble2, David R. Carrier2

1Department of Comparative Biomedical Sciences, Royal Veterinary College, Hatfield, Hertfordshire, United Kingdom, 2Department of Biology, University of Utah, Salt

Lake City, Utah, United States of America

Abstract

Locomotor-respiratory coupling (LRC), phase-locking between breathing and stepping rhythms, occurs in many vertebrates.When quadrupedal mammals gallop, 1:1 stride per breath coupling is necessitated by pronounced mechanical interactionsbetween locomotion and ventilation. Humans show more flexibility in breathing patterns during locomotion, using LRCratios of 2:1, 2.5:1, 3:1, or 4:1 and sometimes no coupling. Previous studies provide conflicting evidence on the mechanicalsignificance of LRC in running humans. Some studies suggest LRC improves breathing efficiency, but others suggest LRC ismechanically insignificant because ‘step-driven flows’ (ventilatory flows attributable to step-induced forces) contribute anegligible fraction of tidal volume. Yet, although step-driven flows are brief, they cause large fluctuations in ventilatory flow.Here we test the hypothesis that running humans use LRC to minimize antagonistic effects of step-driven flows onbreathing. We measured locomotor-ventilatory dynamics in 14 subjects running at a self-selected speed (2.660.1 ms21) andcompared breathing dynamics in their naturally ‘preferred’ and ‘avoided’ entrainment patterns. Step-driven flows occurredat 1-2X step frequency with peak magnitudes of 0.9760.45 Ls21 (mean 6S.D). Step-driven flows varied depending onventilatory state (high versus low lung volume), suggesting state-dependent changes in compliance and damping ofthoraco-abdominal tissues. Subjects naturally preferred LRC patterns that minimized antagonistic interactions and alignedventilatory transitions with assistive phases of the step. Ventilatory transitions initiated in ‘preferred’ phases within the stepcycle occurred 2x faster than those in ‘avoided’ phases. We hypothesize that humans coordinate breathing and locomotionto minimize antagonistic loading of respiratory muscles, reduce work of breathing and minimize rate of fatigue. Future workcould address the potential consequences of locomotor-ventilatory interactions for elite endurance athletes and individualswho are overweight or obese, populations in which respiratory muscle fatigue can be limiting.

Citation: Daley MA, Bramble DM, Carrier DR (2013) Impact Loading and Locomotor-Respiratory Coordination Significantly Influence Breathing Dynamics inRunning Humans. PLoS ONE 8(8): e70752. doi:10.1371/journal.pone.0070752

Editor: Francois Hug, The University of Queensland, Australia

Received March 4, 2013; Accepted June 28, 2013; Published August 12, 2013

Copyright: � 2013 Daley et al. This is an open-access article distributed under the terms of the Creative Commons Attribution License, which permitsunrestricted use, distribution, and reproduction in any medium, provided the original author and source are credited.

Funding: The authors have no support or funding to report.

Competing Interests: The authors declare that co-author David Carrier is a PLOS ONE Editorial Board member. This does not alter the authors’ adherence to allthe PLOS ONE policies on sharing data and materials.

* E-mail: [email protected]

Introduction

Effective ventilation is essential for sustained animal locomotion.

For many animals, including birds and mammals, this requires

integrating movement and breathing so that inspiration and

expiration occur during mechanically compatible periods of the

locomotor cycle. Several direct mechanical links between locomo-

tion and ventilation necessitate integration [1,2,3,4]. Firstly,

sagittal bending of the trunk assists forward progression during

locomotion in quadrupeds, and creates a ‘bellows’ effect, altering

the pressure and volume of the abdomen and thorax. Additionally,

impact loads induce inertial motions of soft-tissues (viscera,

adipose), creating a ‘visceral piston’ effect, pulling and pushing

on the diaphragm and body wall muscles (abdominals, intercostals)

and altering thoraco-abdominal pressures. Finally, many axial

muscles of terrestrial vertebrates contribute to both breathing and

locomotion [5,6,7,8,9,10]. Indeed, many ‘ventilatory muscles’

cease respiratory action and become entrained with the locomotor

cycle during running in lizards [5,6], birds [7] and dogs [8,9]. As a

result of these factors, active inspiration is most compatible with a

specific and different phase of the locomotor cycle than active

expiration.

Locomotor-respiratory coupling (LRC) refers to phase locking

of running and breathing so that the same number of steps occur

during each breath, and has been observed in numerous

vertebrates, including birds, dogs, hares, horses, wallabies and

humans [2,3,11,12,13]. LRC is a form of entrainment, in which

the two rhythmic activities with different frequencies become

phase-locked due to mechanical and neural interactions. LRC has

been suggested to have a number of important physiological

effects. These include reducing the energy cost of breathing

[14,15,16,17], minimizing conflict in muscles that contribute to

both functions [6,9], body stabilization during motion [11,18], and

enabling trunk bending and inertial movements of soft-tissues to

augment pumping of air in and out of the lungs [2,19].

Whereas most galloping mammals exhibit 1:1 (strides/breath)

coupling, humans demonstrate more flexibility in breathing

patterns during locomotion. Humans frequently use LRC ratios

of 2:1, 2.5:1, 3:1, or 4:1 and sometimes lack entrainment

altogether, using independent breathing and stepping frequencies

PLOS ONE | www.plosone.org 1 August 2013 | Volume 8 | Issue 8 | e70752

[2,20,21,22,23]. Variance in LRC pattern is observed both within

and across individuals [14,21,22,23,24]. This flexibility in human

breathing patterns relative to locomotion most likely relates to an

upright posture and striding bipedal gait. Unlike quadrupeds, the

sagittal bending of the back does not assist forward progression

and the forelimbs are not subjected to direct weight-bearing and

impact loading. The forces transmitted through the thorax,

abdomen and rib cage are consequently smaller, reducing

mechanical interactions between locomotion and ventilation.

Ground birds, which are also bipedal, exhibit similar flexibility

as humans in locomotor-ventilatory coupling patterns [3,17].

Bipedal posture may reduce the mechanical and neuromuscular

conflicts that constrain quadrupeds to a 1:1 LRC pattern [8,9].

Does this greater flexibility suggest that breathing and running

are mechanically independent in humans? The answer remains

controversial. Humans do often exhibit sustained ventilatory

entrainment to locomotor forces [2,21,22]. Furthermore, evidence

suggests coupling improves breathing efficiency in some circum-

stances [2,15,24]. On the other hand, humans also couple

breathing frequency with low impact, non-locomotor movements

such as finger and arm tracking [25,26]. Thus, large mechanical

interactions are not required to induce rhythmic coupling.

Theoretically, human LRC during running could arise entirely

from neural interactions.

The overall ventilatory effect of locomotor-ventilatory interac-

tions likely depends on a number of neural and mechanical factors.

The primary mechanical factors in mammalian locomotor-

ventilatory interactions are thoracic loading and inertial displace-

ment of the soft-tissues associated with body accelerations

[2,12,19,20,27,28]. In humans, impact loading of the body at

footstrike is likely to cause compression of the thorax and

downward displacement of the abdominal viscera. Arm swinging

is also likely to exert a compression load on the thorax during early

to mid-stance [29]. These mechanical actions can alter respiratory

flow by changing thoraco-abdominal volumes and pressures.

Inertial displacement of the abdominal viscera would tend to pull

or push on the respiratory diaphragm through its direct

attachment to the liver. The relative magnitude of the effects on

the rib cage and diaphragm likely depends on the mass of the

abdominal organs, the distribution of other soft-tissues, and the

vertical accelerations of the body. Body wall muscles including the

intercostals and abdominals could act to resist these effects [30,31].

Additionally, the rhythmic activity of intercostal and abdominal

muscles is entrained by afferents from the arms and legs,

facilitating neural coupling [32,33]. The ventilatory outcome of

these neuro-mechanical interactions likely depends on the relative

timing and magnitude of the mechanical factors (thorax compres-

sion and visceral displacement) and the rhythmic activation of

body wall muscles.

To address whether locomotor-ventilatory interactions are

mechanically significant in running humans, Banzett and

colleagues [20] measured ‘step-driven flow’ in humans running

on a treadmill. ‘Step-driven flow’ is ventilatory flow and resulting

volumes attributable to locomotor-induced forces, once the

primary ventilatory pattern is removed (through either ensemble

averaging or filtering – see methods). Banzett and colleagues [20]

found the volumes attributable to step-driven flows to be a

negligible fraction of tidal volume, ,1–2% tidal volume per step.

Consequently, they concluded that locomotion has no meaningful

mechanical influence on breathing in humans [20]. Following

publication of this finding, there has been relatively little research

on the mechanical consequences of locomotor-ventilatory inter-

actions in running humans.

Yet, there are reasons to revisit the potential mechanical

significance of locomotor-ventilatory interactions in humans. In

the study by Banzett and colleagues, the subjects lightly rested

their hands on side rails while jogging on the treadmill [20]. Arm

swinging contributes to compression loads on the thorax during

locomotion, and to control of body and head rotation [29,34]. The

restriction of natural arm motion may have reduced mechanical

interactions between locomotion and ventilation. Furthermore,

although the volumes attributed to step-driven flows amounted to

only 1–2% of tidal volume per step, Banzett and colleagues

measured large transient flows associated with footstrike [20].

These step-driven flows lasted only a small fraction of the

ventilatory cycle, but exhibited peak magnitudes of 1 Ls21.

Although brief, step-driven flows reach a large fraction of

concurrent ventilatory flow, sometimes resulting in a brief mid-

breath reversal in flow (from inspiration to expiration or vice-

versa). We suggest that these transient flows have potential to

influence breathing dynamics and ventilatory muscle loading

without directly driving large ventilatory volumes.

Here we test the hypothesis that humans use LRC to minimize

antagonistic effects of step-driven flows on breathing dynamics.

Due the natural variability in coupling patterns in running humans

[14,21,22,23,24], it is possible to compare breathing dynamics

when individuals use their own ‘preferred’ and ‘avoided’

entrainment patterns. We analyze locomotor-ventilatory interac-

tions in fourteen physically fit subjects running on a treadmill at a

moderate self-selected speed (2.660.1 ms21) with natural arm

motion. Self-selected speed is likely to be close to energetically

optimal speed [35]. Running at speeds above preferred speeds

would increase both impact loads and ventilatory demand,

increasing the potential for antagonistic locomotor-ventilatory

interactions. We therefore consider self-selected speed to provide a

conservative test of the hypothesis that humans use LRC to

minimize antagonistic locomotor-ventilatory dynamics.

Materials and Methods

Ethics StatementThe Institutional Review Board of University of Utah approved

the protocols and informed consent documents, under IRB

#12361, and all subjects gave written informed consent.

Subjects and protocolFourteen adult subjects volunteered for this study (9 females 5

males; age 3662 years, mass 65.162.8 kg, height 1.7260.02 m).

Seven of the subjects were recreational runners who regularly ran

at least 20 miles per week, and seven were physically fit non-

runners (Table 1). This sampling was designed to obtain a broad

representation of human locomotor-ventilatory patterns, not to

statistically compare trained and non-trained runners, which has

been previously studied by others [23,24]. Participants ran at a

steady speed on a treadmill, using natural arm motion. After an

initial practice session, subjects were allowed to select their

preferred speed, with the instructions to run at a comfortable pace

that they could sustain for 30 minutes. They warmed up for

5 minutes and then continued to run for 5 minutes while we

recorded running and breathing patterns.

Data collectionWe tracked the step cycle by measuring acceleration at the top

of the head using an Endevco microtron accelerometer (model

7290A-10). The subjects wore a lightweight facemask equipped

with two-way non-rebreathing valves to measure inspiratory and

expiratory flow separately (Hans Rudoph Inc., Kansas City, MO,

Locomotor-Ventilatory Dynamics in Humans


USA). Each valve was connected by tubing (34.2 mm ID; Hans

Rudolph, Inc.) to screen pneumotachs (model 4813, Hans

Rudolph Inc.), with a calibrated flow range of 0–800 liters per

minute. The pressure ports of the screen pneumotachs were

attached to a differential pressure transducer (Omega Engineering,

Stamford CT, USA). To prevent motion artifact, the tubing and

pneumotachs were secured to a small backpack worn by the

subject, fitted with both waist and shoulder straps. Flow was

calibrated by recording the output voltage from a known volume

of air, using a 3 liter calibration syringe (Hans Rudolph Inc.).

The signals were amplified 10x (model P122, Grass Inc., West

Warwick RI, USA), sampled at 200 Hz and digitally recorded

onto computer using the MP100 data acquisition system with

AcqKnowledge software (BIOPAC Systems Inc., Goleta, CA,

USA.). Before analysis, data were low pass filtered at 30 Hz using

a zero-phase 6th order Butterworth filter in Matlab (butter and

filtfilt functions; Matlab release 13). These filter settings were

selected because they were found to effectively remove noise while

retaining higher frequency signal characteristics such as peaks

associated with footstrike impact.

Calculation of average step-driven flowTo enable comparison to previous findings, we calculate

average step-driven flow using similar analytical methods of

Banzett and colleagues [20]. We calculated an ensemble average

(EA) of ventilatory flow relative to the step cycle for each

individual, including at least 50 steps per subject [20]. The EA

technique samples flow segments distributed equally through the

ventilatory cycle, to quantify step-driven flow independent of

ventilatory phase. We calculated this EA for the 12 out of 14

subjects who exhibited varying phase of entrainment between

running and breathing cycles (Table 1). The resulting EA trace of

step-driven flow was integrated over time to obtain the locomotor

driven volume (LDV). Expiratory (positive) and inspiratory

(negative) volumes were integrated separately to assess the relative

magnitude of each. The LDV was normalized by the total

concurrent ventilatory volume (Vtot) over the same time periods as

the EA flow segments, to obtain a value as a percentage of

concurrent ventilatory flow. Note that our calculation of LDV

differs slightly from that used by Banzett and colleagues, who

reported the fractional contribution to tidal volume per step;

whereas we report the total fractional contribution to ventilation.

Humans typically use coupling ratios of 2:1 and higher (e.g.

[2,20,21,22,23]), taking multiple steps in each breath cycle;

consequently, a per step value does not provide a complete account

of the total magnitude of locomotor-driven ventilation.

A potential source of error in the calculated LDV is flow artifact

caused by motion of the tubes during running. As mentioned

earlier, we minimized this by securing the tubes to a lightweight

backpack worn by the subject. To quantify the error associated

with any remaining motion, we calculated an ‘artifact LDV’ for 5

subjects. We recorded inspiration and expiration in separate

channels, causing each channel to have a baseline of zero during

the opposing ventilatory half-cycle. During this period, the only

flow measured in that channel was small magnitude, high

frequency artifact flow due to tube motion. To quantify this, we

calculated a value of LDV using the same method described

above, except that the flow segments in the ensemble average were

obtained from the baseline of the inactive ventilation channel. This

method resulted in an artifact LDV value of 0.760.1% of Vtot

(mean 6 s.e.m.) during control running trials, with a worst case

artifact magnitude of 1% Vtot.

Testing for variation in step-driven flow with ventilatoryphaseThe visceral piston hypothesis predicts that the effect of

locomotor forces on flow depends on the phase of the breath

cycle in which the step occurs [2] (Fig. 1C). Deceleration of the

body at footstrike likely has two primary mechanical effects: 1)

compression of the thorax, and 2) viscoelastic bounce and rebound

of the guts [20,28]. Contraction of respiratory muscles and

inflation of the lungs differentially alters thoracic and abdominal

Table 1. Subject data.

Subject Gender Age (yrs) Mass (kg) Height (m) Speed (ms) Training Entrainment Coupling Ratio (s)

1 F 39 54.0 1.57 2.4 Casual Coupled, vp,vr* 2:1, 3:1

2 F 41 59.4 1.73 2.4 Nonrunner Coupled 2:1

3 M 35 72.5 1.75 2.9 Nonrunner Coupled, vr* 3:1, 3.5:1

4 M 43 76.2 1.83 2.4 Nonrunner Uncoupled*

5 F 43 60.8 1.73 2.6 Casual Coupled, vp* 2:1

6 F 23 59.0 1.65 2.4 Nonrunner Coupled, vp* 2:1

7 M 26 86.2 1.78 2.4 Nonrunner Coupled, vp,vr* 2:1, 2.5:1, 3:1

8 F 41 51.9 1.67 2.4 Casual Coupled, vp* 2.5:1

9 F 28 64.4 1.70 2.4 Nonrunner Uncoupled*

10 M 49 71.7 1.78 2.8 Trained Coupled, vr * 2:1, 3.5:1

11 F 47 62.6 1.65 2.4 Trained Coupled, vp,vr* 2:1,3:1

12 F 27 59.0 1.75 3.4 Trained Coupled,vp* 2:1

13 F 25 54.5 1.60 3.4 Trained Coupled 2:1

14 M 31 79.4 1.83 2.3 Nonrunner Uncoupled*

mean 36 65.1 1.72 2.6

s.e.m. 2 2.8 0.02 0.1

*Ensemble average calculated for this subject.Coupling abbreviations: vr = variable coupling ratio. vp = variable coupling phase relationship.doi:10.1371/journal.pone.0070752.t001



stiffness [36]. Locomotor forces are likely to have an enhanced

expiratory effect if they occur when the abdominal muscles are

being recruited for forced expiration, because these muscles resist

the descent of the guts (Fig. 1C).

The ensemble average technique cannot reveal this effect of

ventilatory phase because it averages step-driven flow across all

phases. Therefore, we also compared step-driven ventilation at

different phases of the breathing cycle, in bins of 20% of maximum

tidal volume (Vmax) during inspiration and expiration. From the

ensemble averaged data, we observed that step related oscillations

in flow occurred at step frequency and the 1st and 2nd harmonic of

step frequency. We high-pass filtered the flow signal to remove the

primary breathing pattern, so that only flows associated with the

step frequency or higher remained in the signal. We used a zero-

phase digital Butterworth filter (‘butter’ and ‘filtfilt’ functions in

Matlab version 6.5), with the cutoff frequency and filter order

specified using the Matlab function ‘buttord’ to create a filter that

lost no more than 1% of the signal greater than or equal to step

frequency and attenuated 99% of the signal below half the step

frequency. We then calculated an average 6 s.e.m of step-driven

flow for all steps occurring within each volume bin of the

ventilatory cycle.

Analysis of entrainment patternsWe used phase analysis to examine the locomotor-ventilatory

coupling patterns. The timing of each inspiratory and expiratory

transition was calculated relative to the step cycle. The beginning

of footstrike was defined as zero, and the phase of each ventilatory

transition was expressed in degrees between 0 and 360 or as a

fraction of the step cycle by dividing the phase angle by 360. The

step cycle was divided into 20 bins (18 degrees or 5% of the step

cycle each), and the frequency of inspiratory and expiratory

transitions was calculated for each bin. If breaths and steps occur

randomly with respect to each other, the distribution of ventilatory

transition events should not significantly differ from a uniform

circular distribution. For each subject whose ventilatory transitions

significantly differed from a uniform circular distribution (see

Statistics), we determined the ‘preferred’ phase for inspiratory and

expiratory initiation as the bin in which each of these events

occurred most frequently. Likewise, we determined ‘avoided’

phase as the bin in which each transition event occurred least

frequently.

Ventilatory transition timesTo quantify the time required for transition between ventilatory

half cycles (inspiration and expiration), we measured the time

between 50% peak flows (T50). The expiratory T50 is the time

from 50% peak inspiration to 50% peak expiration. Likewise, the

inspiratory T50 is the same measure between 50% peak expiration

and 50% peak inspiration. We compared T50 values between

‘preferred’ and ‘avoided’ phase bins, described above. If multiple

bins were tied for ‘preferred’ or ‘avoided’, we took the average of

the tied bins.

StatisticsKuiper’s test of circular uniformity was used to test whether

inspiratory and expiratory initiation events were distributed

randomly with respect to the step cycle. For subjects with

ventilatory transitions distributed non-uniformly relative to the

step cycle, a Friedman nonparametric repeated measures ANOVA

test was used to compare ventilatory transition times (T50)

between preferred and avoided phases of step cycle, with posthoc

pairwise comparisons using Dunn’s Multiple Comparisons.

Results

Step-driven oscillations in flowAll subjects exhibited high frequency, step-driven oscillations

in flow (Fig. 1B). An inspiratory pulse occurred immediately after

footstrike, followed by an expiratory pulse as the body

approached peak acceleration (Fig. 2). The associated inspiratory

and expiratory volumes amounted to 212.764.5% (mean 6S.D)

and 10.763.2%, respectively, of the concurrent ventilatory

volume (Fig. 2; Table 2), with peak flow magnitude averaging

0.9760.45 Ls21 across individuals. In some cases, these step-

driven flows were large enough to cause transient reversals in

flow (Fig. 3A).

Locomotor-ventilatory interactions and entrainmentAlthough the step-driven inspiratory and expiratory volumes

were equal on average (Fig. 2), the effect of locomotor acceleration

Figure 1. Locomotor-ventilatory interactions. (A–B) Typical head acceleration (top) and ventilatory flow (bottom: expiration positive,inspiration negative) during quiet standing (A), and moderate speed treadmill running (B). Note the high frequency oscillations in ventilatory flowduring running. (C) Schematic illustration of the ‘visceral-piston’ model for human locomotor-ventilatory interactions. Red arrows indicate muscleactions during inspiration and expiration.doi:10.1371/journal.pone.0070752.g001



Figure 2. Step-driven ventilatory flows and volumes. (A) Grand mean of step-driven ventilatory flow across subjects (mean6 95% confidenceinterval for 12 subjects, see also individual examples in Fig. 5). Expiration is positive. (B) Mean 6 SD of step-driven volume as a percentage of totalconcurrent ventilatory volume (Vtot), during level moderate speed running (N= 12, black dots show data from individuals). Step-driven flow andvolume data included here are from subjects with variation in the phase locking between steps and breaths (Table 1).doi:10.1371/journal.pone.0070752.g002

Figure 3. Step driven flow depends on ventilatory phase. (A) Variation in the magnitude of step-driven flows is apparent, particularly duringlow frequency breathing, as shown here during 3:1 (strides per breath) coupling. Note the reversal in flow in late expiration (asterisk). Dashed verticallines indicate footstrike events (data from subject 3). (B) Average step-driven flow (means 695%CI) for a representative subjective at four points inthe ventilatory cycle: 1) early expiration, high lung volume (90% Vmax), 2) late expiration, low lung volume (10% Vmax), 3) early inspiration, low lungvolume (10% Vmax) and 4) late inspiration, high lung volume (90% Vmax).doi:10.1371/journal.pone.0070752.g003



on flow varied considerably with ventilatory phase (Fig. 3). Across

individuals, we observed a significant relationship between tidal

volume at footstrike and the net step-driven volume (Fig. 4). Steps

occurring in late expiration (low lung volume) tend to have a net

inspiratory effect. Similarly, steps occurring in late inspiration

(high lung volume) tend to have a net expiratory effect. Thus, the

net effect of step-driven flow is synergistic early in the ventilatory

half-cycle and antagonistic late in each ventilatory half cycle. Flow

reversals occur most often in late expiration during slow, deep

breaths and coupling with ratios greater than 2.5:1 strides per

breath (Fig. 3A). However, most runners avoided these flow

reversals because they breathed more frequently, closer to a 2:1

stride per breath rhythm.

Runners varied in whether or how tightly they entrained

breathing and stepping rhythm (Table 1). Most (11 of 14) subjects

exhibited periods of fixed-ratio coupling between stepping and

breathing rhythms, with 2:1 strides per breath as the most

common pattern. Two of these subjects exhibited strict phase-

locked coupling with a single coupling ratio, whereas the

remaining 9 subjects exhibited varied phase or switching among

multiple coupling ratios (Table 1). Despite variation in coupling,

13 of 14 subjects had significantly non-uniform distributions of

ventilatory transitions relative to the step cycle (inspiratory,

expiratory or both; Table 3). Preferred transition phases tended

to correspond to regions of the step cycle that mechanically

assisted ventilation, or at least, did not impede it (Fig. 5).

Although step-driven flows are brief, the phasing of transitions

relative to the step cycle does substantially influence breathing

dynamics. Whether coupled or not, most runners avoided reversals

in flow by using a breathing frequency near a 2:1 stride per breath

ratio and timing ventilatory transitions to coincide favorably with

step-driven flow (Fig. 5). Furthermore, the phasing between steps

and breath transitions influenced the time required for ventilatory

transitions. When runners coordinate the transitions to occur in

phases of the step cycle that promote flow, transitions occur 2X

more rapidly (Fig. 6). When compared across all uncoupled or

variably coupled runners (for which ‘avoided’ transition data were

available), ventilatory transitions in ‘preferred’ phase relationships

with the step cycle occurred more rapidly than those in ‘avoided’

phases (Fig. 6). The difference between preferred and avoided

transition T50 values was statistically significant (p = 0.0002 for

Freidman Test, p,0.05 Dunn’s Multiple Comparisons).

Discussion

The goal of this study was to examine whether the mechanical

interaction between running and breathing in humans is large

enough to be physiologically important. We measured the

ventilatory flows and volumes attributable to step-driven flows

during running at a moderate self-selected speed. The timing of

step-driven flows is consistent with the visceral piston hypothesis

for locomotor-ventilatory interactions in humans [2,19,20,27,28].

Table 2. Step-driven ventilatory volume, as a fraction ofconcurrent volume, during inspiration and expiration.

Subject Insp Exp

1 217.2% 15.2%

3 29.8% 9.3%

4 27.7% 7.6%

5 210.5% 7.5%

6 29.6% 9.2%

7 211.5% 8.2%

8 29.9% 8.1%

9 210.1% 9.6%

10 214.7% 14.8%

11 218.8% 14.2%

12 222.2% 15.4%

14 29.9% 8.9%

mean 212.7% 10.7%

s.d. 4.5% 3.2%

doi:10.1371/journal.pone.0070752.t002

Figure 4. The net effect of step-driven ventilation shifts from synergistic to antagonistic in each ventilatory half-cycle. We show netstep-driven volumes (in milliliters per step) as a function of the tidal volume at the time of footstrike during inspiration (left) and expiration (right),averaged across individuals (mean 6 95% CI, N = 12). At low lung volumes, the inspiratory pulse at footstrike is larger, leading to a net inspiratoryeffect. At high lung volumes, a larger expiratory pulse occurs, leading to a net expiratory effect.doi:10.1371/journal.pone.0070752.g004



Step-driven ventilation averaged around 10–12% of total venti-

latory volume with flow magnitudes around 1 Ls21. We found that

step-driven flow dynamics varied depending on ventilatory state

(high versus low lung volume), suggesting phase-dependent changes

in compliance and damping of thoraco-abdominal tissues. We also

discovered that the timing of impact loading relative to the

ventilatory cycle significantly influenced breathing dynamics.

Ventilatory transitions initiated in preferred (assistive) phases of

the step cycle occurred 2x faster than those in avoided

(antagonistic) phases. These findings suggest a physiologically

significant mechanical interaction exists between locomotion and

ventilation in humans.

The step-driven ventilatory volumes we measured appear to be

larger on average than those reported previously [20]. Here, step-

driven volumes amounted to 10–12% of total ventilation, which

amounts to 2.5–3.0% tidal volume per step during a 2:1 step per

breath rhythm. Banzett and colleagues reported a value of 1–2%

tidal volume per step. Nonetheless, the timing of peak step-driven

flow was consistent with Banzett and colleagues. The difference in

magnitude could be explained by a difference in arm motion

between the two studies. In the earlier study, the subjects lightly

rested their hands on side rails, whereas in the current study the

subjects were allowed to move their arms naturally while running.

Both step-induced thoracic loading and inertial displacement of

abdominal viscera likely contribute to locomotor-ventilatory

interactions while running [2,12,19,20,27]. Minimizing the effects

of arm loading on thoracic compression during running could

reduce both the mechanical and neural interactions between

locomotion and ventilation [29,32,33,37]. Nonetheless, we do find

a similar pattern, and although the total step-driven volumes are

small on a per step basis, they do have a significant mechanical

influence on breathing dynamics.

What is the physiological significance of locomotor-ventilatory entrainment in humans?The frequency of step related flows is too high relative to breath

frequency to allow direct coupling to drive ventilation (Figs. 1, 2).

Yet, our data reveal that runners prefer to time ventilatory

transitions to periods in the step cycle that assist rather than

impede flow (Fig. 5), and this timing significantly reduces the time

required for transitions (Fig. 6). These data suggest that step-driven

flows have potential to influence the work of respiratory muscles,

because they significantly influence breathing dynamics, particu-

larly at ventilatory transitions (expiration to inspiration and vice

versa). Appropriate coordination of stepping and breathing rhythm

may act to minimize antagonistic loading of the respiratory

muscles caused by motions of the abdomen and chest wall.

Although locomotor-ventilatory coupling is more flexible in

humans than in quadrupedal mammals [2,3,11,12,13,22], cou-

pling in humans may reduce conflicting demands placed on the

Figure 5. Subjects prefer to initiate ventilatory transitions atphases that assist rather than imped flow. The distribution ofventilatory transitions relative to the step cycle is non-uniform (Table 3).Here we show the net bias in ventilatory transitions relative to stepcycle (left axis and bars), with transitions to expiration as positive,transition to inspiration as negative. The ‘net bias’ value is the numberof expiratory transitions minus the number of inspiratory transitions ineach phase bin. Step-driven flow is overlaid for reference (right axis andlines). Data from 3 individuals illustrates typical variation betweenstrongly coupled (A), variably coupled (B) and uncoupled (C) subjects.Although variation exists, the timing of transitions is clearly non-random, and exhibits some correspondence to the step-driven flow.doi:10.1371/journal.pone.0070752.g005

Table 3. Mean phase and dispersion (in degrees) ofventilatory transitions relative to the step cycle.

Phase relative to step cycle (degrees)

Subject Inspiration Expiration

N mean disp mean disp

1 57 296* 13.0 121* 6.2

2 72 32* 0.1 352* 0.2

3 48 345* 0.4 285* 0.6

{4 70 286 16.0 282 11.6

5 86 176* 0.9 164* 3.0

6 82 28* 0.5 313* 1.2

7 62 297* 1.1 273* 9.5

8 87 325* 0.8 255* 1.1

9 55 203 99.0 334* 0.9

10 35 296* 0.5 356* 0.5

11 79 153* 6.2 154* 1.8

12 90 282* 1.2 150* 8.2

13 87 308* 0.4 212* 0.4

14 62 324 26.0 17* 13.6

*Indicates significant difference from uniform circular distribution based onKuiper’s test.{Indicates subject with no statistical evidence of transition entrainment ininspiration or expiration.doi:10.1371/journal.pone.0070752.t003



diaphragm and body wall muscles (abdominal muscles and

intercostals).

The 2:1 LRC pattern preferred by human runners may also

reflect optimization to minimize antagonistic loading of respiratory

muscles. Humans prefer a 2:1 ratio across a wide range of

sustainable running speeds (e.g. [2,20,21,22,23]). A recent study by

O’Holloran and colleagues found the highest flexibility in LRC

patterns at preferred stride frequency, when energy cost is

minimal. At stride frequencies above or below preferred, energy

demand sharply increases and subjects exhibit a stronger

preference for 2:1 coupling [22]. Here, we find evidence that a

2:1 pattern could minimize work of respiratory muscles by

allowing one footstrike to assist the ventilatory transition and the

second to occur at intermediate lung volumes (see Fig. 6A). At

intermediate lung volumes, the net effect of step-driven flows is

either assistive or neutral (Fig. 4). However, near the end of each

ventilatory half-cycle, approaching the extrema of lung volume,

step-driven flows become antagonistic (Fig. 4). Higher coupling

ratios require footstrikes near the end of each ventilatory half-

cycle, leading to antagonistic loading of the respiratory system and

transient reversals in flow (Fig. 3). Thus, a 2:1 stride per breath

rhythm might be more strongly preferred during intense

endurance running, when fatigue of respiratory muscles could be

limiting.

We hypothesize that human runners benefit from locomotor-

ventilatory entrainment by reducing the work of ventilatory

muscles, and minimizing fatigue of respiratory muscles that are

critical to endurance aerobic activity. Appropriate tuning of

locomotor-ventilatory interactions likely minimizes antagonistic

loading of ventilatory muscles and allows inertial displacement of

the guts to passively assist the action of respiratory muscles. These

mechanical interactions likely have the greatest impact on

diaphragm performance because the abdominal viscera directly

attach to this muscle. Work of breathing increases as a squared

function of ventilatory demand, and breathing may account for up

to 10–15% of energy demand in intense exercise [38,39].

Although the properties of the mammalian diaphragm appear to

confer considerable resistance to fatigue [40], some evidence

suggests that it can be vulnerable to fatigue during prolonged or

intense exercise in both sedentary and fit individuals

[41,42,43,44,45,46,47]. Declines in respiratory function following

marathon and ultra-marathon competitions also suggest respira-

tory muscle fatigue [48,49,50]. Respiratory muscle fatigue may be

a limiting factor in human endurance activity, and locomotor-

ventilatory coupling has potential to minimize fatigue, especially

during activities that involve impact loading with each footstrike,

such as walking and running.

Unfortunately, it remains challenging to directly test this

hypothesis. Ventilatory muscles are a relatively small fraction of

the metabolically active tissue in the body, making it difficult to

measure changes in respiratory muscle work using standard

respirometry techniques. Though technically challenging, it may

be possible to use recordings of muscle activity (e.g., [51]) to

examine the response of ventilatory muscles to manipulations of

visceral load and locomotor-ventilatory coordination during

running.

In future work, it would be interesting to investigate whether

antagonistic locomotor-ventilatory interactions could explain, in

part, why obese individuals experience ‘breathlessness’ and rapid

fatigue during locomotion. Obese individuals face a dual problem

of increased energy cost and impaired respiratory function during

locomotion. Obese individuals incur a 10–25% higher metabolic

energy cost of walking per kilogram body mass compared to

people with healthy weight [52,53]; meaning that carrying fat costs

more than carrying lean weight. The source of this added cost

remains controversial. External work does not explain the added

cost, because the gait of obese individuals involves similar total

mechanical work on their body center of mass [54]. Increased

internal work associated bouncing soft tissues and increased costs

for stabilizing the body and joints may contribute to increased

locomotor costs due to excess fat mass [53,54,55,56,57,58].

Figure 6. Phasing of steps relative to breaths has a significanteffect on the duration of ventilatory transitions. (A) Examples ofventilatory transitions timed with preferred and assistive phases of thestep cycle, facilitating rapid transitions, and (B) timed with an avoidedand antagonistic phases of the step cycle. Dashed vertical lines indicatezero-crossings of ventilatory transitions. The transition durations (T50)are calculated between the black dots indicating the times of 50% peakflows. (C) Transition times averaged across individuals (mean 6 s.e.m),comparing breaths at ‘avoided’ and ‘preferred’ phase relationships.Preferred refers to the most used phase bin (e.g., see Fig. 5), and‘avoided’ refers to the least used phase bin that was represented in thedata.doi:10.1371/journal.pone.0070752.g006



Furthermore, excess fat is associated with reduced operating lung

volumes and decreased respiratory compliance, factors that

increase the work of breathing [59,60]. Soft-tissue bouncing

induced by impact loads during walking and running may further

increase the work of breathing and the rate of respiratory muscle

fatigue in overweight and obese individuals. Whereas in lean

individuals a large fraction of soft-tissue mass is concentrated intra-

abdominally, obesity results in a large fraction of soft-tissue mass

distributed external to the body wall muscles. Thus, high adiposity

will increase soft-tissue bouncing and reduce the ability to actively

tune soft-tissue dynamics through abdominal muscle contraction.

Larger, uncontrolled soft-tissue motions may result in antagonistic

locomotor-ventilatory interactions and higher mechanical work of

breathing during exercise. We predict that the problems of obesity

are exacerbated by antagonistic locomotor-ventilatory interactions

during walking and running.

ConclusionsBiomechanics studies often focus on musculoskeletal and

biomechanical factors as limits to performance. For example, leg

muscle strength is widely thought to limit top running speed

during sprinting [61,62]. However, during endurance locomotion,

respiratory muscles, not leg muscles, may limit maximum exercise

intensity and duration [47]. We suggest that when assessing the

physiological importance of locomotor-respiratory coupling in

humans, it may be short sighted to place too much emphasis on

the comparatively small volumes attributed to it. Entrainment

might serve a number of beneficial physiological functions – i.e.,

reducing the work of ventilatory muscles, preventing respiratory

muscle fatigue, and improving respiratory efficiency through

enhanced gas mixing, transport and exchange. Consequently, the

precise volume driven by locomotion might be less important than

the associated intrapulmonary dynamics. Humans are well

adapted for economic walking and endurance running

[34,63,64,65]. Multiple derived characters suggest aerobic endur-

ance as a key evolutionary pressure in the human lineage,

including skeletal morphology, developed tendon springs and

enhanced heat dissipation [34,63,64,66]. We will not be surprised,

therefore, if the unusual locomotor-respiratory coupling patterns

exhibited by human runners proves to be another example of

evolutionary adaptation in support of the exceptional endurance

running capabilities of humans [63,64].

Acknowledgments

We thank Eric Stakebake for assistance in the experiments, Franz Goller

for loan of equipment and the late Farish Jenkins Jr. for helping to make

the project possible. We also thank Joanne Gordon for her comments on

the manuscript draft.

Author Contributions

Conceived and designed the experiments: MAD DRC DMB. Performed

the experiments: MAD. Analyzed the data: MAD. Wrote the paper: MAD.

Interpreted the data and revised the manuscript: DMB DRC.

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